Porous biodegradable polymeric materials for cell transplantation

ABSTRACT

Polymeric materials are used to make a pliable, non-toxic, injectable porous template for vascular ingrowth. The pore size, usually between approximately 100 and 300 microns, allows vascular and connective tissue ingrowth throughout approximately 10 to 90% of the matrix following implantation, and the injection of cells uniformly throughout the implanted matrix without damage to the cells or patient. The introduced cells attach to the connective tissue within the matrix and are fed by the blood vessels. The preferred material for forming the matrix or support structure is a biocompatible synthetic polymer which degrades in a controlled manner by hydrolysis into harmless metabolites, for example, polyglycolic acid, polylactic acid, polyorthoester, polyanhydride, or copolymers thereof. The rate of tissue ingrowth increases as the porosity and/or the pore size of the implanted devices increases. The time required for the tissue to fill the device depends on the polymer crystallinity and is less for amorphous polymers versus semicrystalline polymers. The vascularity of the advancing tissue is consistent with time and independent of the biomaterial composition and morphology.

BACKGROUND OF THE INVENTION

This is a continuation of continuation of U.S. Ser. No. 08/669,760,filed Sep. 26, 2000, which is a continuation of U.S. Ser. No.08/052,387, filed Apr. 23, 1993, now abandoned, which is acontinuation-in-part of U.S. Ser. No. 08/012,270, filed Feb. 1, 1993,now issued as U.S. Pat. No. 5,514,378.

This invention is generally in the field of polymeric materials, and inparticular in the area of biocompatible artificial matrices forimplantation of cells.

Loss of organ function can result from congenital defects, injury ordisease. Many times treatment with drugs or surgery is not in itselfsufficient and the patient dies or is severely disabled. One approachfor treatment has been to transplant donor organs or tissue into thepatient. Drugs such as cyclosporin can be used to prevent tissuerejection. However, there is a tremendous shortage of donor organs, mostof which must come from a recently deceased individual. There have beena number of attempts to culture dissociated tissue and implant the cellsdirectly into the body. One of the problems with implanting dissociatedcells into the body is that they do not form three dimensionalstructures and the cells are lost by phagocytosis and attrition. Oneapproach to overcome this problem is described by U.S. Pat. No.4,352,883 to Lim, wherein cells are encapsulated within alginatemicrospheres, then implanted. While this method can sometimes maintainviable functioning cells, the cells do not form organs or structures andrarely result in long term survival and replication of the encapsulatedcells. Most cells have a requirement for attachment to a surface inorder to replicate and to function.

The first attempts to culture cells on a matrix for use as artificialskin, which requires formation of a thin three dimensional structure,were described by Yannas and Bell in a series of publications. They usedcollagen type structures which were seeded with cells, then placed overthe denuded area. A problem with the use of the collagen matrices wasthat the rate of degradation is not well controlled. Another problem wasthat cells implanted into the interior of thick pieces of the collagenmatrix failed to survive.

One method for forming artificial skin by seeding a fibrous lattice withepidermal cells is described in U.S. Pat. No. 4,485,097 to Bell, whichdiscloses a hydrated collagen lattice that, in combination withcontractile agents such as platelets and fibroblasts and cells such askeratinocytes, is used to produce a skin-equivalent. U.S. Pat. No.4,060,081 to Yannas et al. discloses a multilayer membrane useful assynthetic skin which is formed from an insoluble non-immunogenicmaterial which is nondegradable in the presence of body fluids andenzymes, such as cross-linked composites of collagen and amucopolysaccharide, overlaid with a non-toxic material such as asynthetic polymer for controlling the moisture flux of the overallmembrane. U.S. Pat. No. 4,458,678 to Yannas et al. discloses a processfor making a skin-equivalent material wherein a fibrous lattice formedfrom collagen cross-linked with glycosaminoglycan is seeded withepidermal cells. A disadvantage to the first two materials is that thematrix is formed of a “permanent” synthetic polymer. In the third case,the matrix can be biodegradable but, since it is formed primarily ofcollagen, only by enzymatic action, which occurs in an uncontrolledmanner.

U.S. Pat. No. 4,520,821 to Schmidt describes a similar approach that wasused to make linings to repair defects in the urinary tract. Epithelialcells were implanted onto synthetic non-woven biodegradable polymericmatrices, where they formed a new tubular lining as the matrix degraded.The matrix served a two fold purpose—to retain liquid while the cellsreplicated, and to hold and guide the cells as they replicated. However,this approach is clearly limited to repair or replacement of very thinlinings.

Vacanti, et al., Arch. Surg. 123:545-549 (1988), describe a method ofculturing dissociated cells on biocompatible, biodegradable matrices forsubsequent implantation into the body. This method was designed toovercome a major problem with previous attempts to culture cells to formthree dimensional structures having a diameter of greater than that ofskin. Vacanti and Langer recognized that there was a need to have twoelements in any matrix used to form organs: adequate structure andsurface area to implant a large volume of cells into the body to replacelost function and a matrix formed in a way that allowed adequatediffusion of gases and nutrients throughout the matrix as the cellsattached and grew to maintain viability in the absence ofvascularization. Once implanted and vascularized, the porosity requiredfor diffusion of the nutrients and gases was no longer critical.

However, even with the method described by Vacanti, the implant wasinitially constructed in vitro, then implanted. It is clearly desirableto be able to avoid the in vitro step. It is also desirable to havebetter ways that can be used to form synthetic, biodegradable matricesthat can be implanted and sustain cell growth in vivo, degrading in acontrolled manner to leave functional, viable cells organized to form anorgan equivalent.

It is therefore an object of the present invention to provide apolymeric material which can be implanted into the body, vascularizedand used as a means to achieve a high survival rate for dissociatedcells injected into the matrix.

It is a further object of the present invention to provide abiocompatible, polymeric implant which can be implanted with cellswithout prior in vitro culturing and then degrades at a controlled rateover a period of time as the implanted cells replicate and form an organstructure.

SUMMARY OF THE INVENTION

Polymeric materials are used to make a pliable, non-toxic, implantableporous template for vascular ingrowth and into which cells can beinjected. The pore size, usually between approximately 100 and 300microns, allows vascular and connective tissue ingrowth throughoutapproximately 10 to 90% of the matrix following implantation, and theinjection of cells uniformly throughout the implanted matrix withoutdamage to the cells or patient. The introduced cells attach to theconnective tissue within the matrix and are fed by the blood vessels.The preferred material for forming the matrix or support structure is abiocompatible synthetic polymer which degrades in a controlled manner byhydrolysis into harmless metabolites, for example, polyglycolic acid,polylactic acid, polyorthoester, polyanhydride, or copolymers thereof.The elements of these materials can be overlaid with a second materialto enhance cell attachment. The polymer matrix is configured to provideaccess to ingrowing tissues to form adequate sites for attachment of therequired number of cells for viability and function and to allowvascularization and diffusion of nutrients to maintain the cellsinitially implanted.

As described in the examples, highly-porous biocompatible andbiodegradable polymers forms were prepared and implanted in themesentery of rats for a period of 35 days to study the dynamics oftissue ingrowth and the extent of tissue vascularity, and to exploretheir potential use as substrates for cell transplantation. Theadvancing fibrovascular tissue was characterized from histologicalsections of harvested devices by image analysis techniques. The rate oftissue ingrowth increased as the porosity and/or the pore size of theimplanted devices increased. The time required for the tissue to fillthe device depended on the polymer crystallinity and was less foramorphous polymers versus semicrystalline polymers. The vascularity ofthe advancing tissue was consistent with time and independent of thebiomaterial composition and morphology.

BRIEF DESCRIPTION OF THE DRAWINGS

FIGS. 1 a, b, c, and d are digitized images of cross sections ofsemicrystalline PLLA (L90c) transplantation devices harvested at 5, 15,25, and 35 days. The shadowed area defines the region which waspartially filled with ingrowing tissue.

FIGS. 2 a, 2 b, 2 c, and 2 d, are graphs of the normalized tissueingrowth (FIG. 2 a), fraction of tissue area of prevascularized regions(FIG. 2 b), number of capillaries per field (FIG. 2 c), and averagecapillary area (μm²) (FIG. 2 d), as a function of implantation time(days), for prewet semicrystalline PLLA (L90e) devices implanted in thedistal (n) and proximal site (1) of the mesentery. The error bars forthe tissue ingrowth designate averages±range of two experiments whereasthose for the tissue area, the number of capillaries, and the averagecapillary area stand for averages±standard deviation of three 1×1 mm²fields of the same histological section.

FIGS. 3 a, 3 b, 3 c, and 3 d, are graphs of the 5 normalized tissueingrowth (FIG. 3 a), fraction of tissue area of prevascularized regions(FIG. 3 b), number of capillaries per field (FIG. 3 c), and averagecapillary area (μm²) (FIG. 3 d), as a function of implantation time(days), for semicrystalline PLLA devices of different porosities andpore sizes [L90e (squares) L90c (circles); and L80e (triangles)]implanted in a dry form. The error bars for the tissue ingrowthdesignate averages±range of two experiments whereas those for the tissuearea, the number of capillaries, and the average capillary area standfor averages±standard deviation of three 1×1 mm² fields of the samehistological section.

FIGS. 4 a, 4 b, 4 c, and 4 d, are graphs of the normalized tissueingrowth (FIG. 4 a), fraction of tissue area of prevascularized regions(FIG. 4 b), number of capillaries per field (FIG. 4 c), and averagecapillary area (μm²) (FIG. 4 d), as a function of implantation time(days), for amorphous PLLA. devices of different pore sizes [NCL90e (n)and NCL90c (1)] implanted in a dry form. The error bars for the tissueingrowth designate averages±range of two experiments whereas those forthe tissue area, the number of capillaries, and the average capillaryarea stand for averages±standard deviation of three 1×i mm² fields ofthe same histological section.

FIG. 5 is a graph of the percent normalized tissue ingrowth as afunction of implantation time (days) for PLGA (85:15) (85LG90e) (l) andPLGA (50:50) (50LG90e) (s) devices implanted in a dry form. The errorbars designate averages±range of two experiments.

DETAILED DESCRIPTION OF THE INVENTION

As described in more detail below, the present invention is thepreparation and use of synthetic, biocompatible, biodegradable polymericmatrices for implantation into a patient, followed by seeding of cells.In the preferred method, the matrix is implanted, vascularized byingrowth of capillaries and connective tissue from the recipient, thenthe cells are seeded.

As demonstrated by the examples, the preferred matrix is an amorphous orsemicrystalline polymer such as poly(lactic acid-glycolic acid) having aporosity (defined herein as the fraction of void volume) in the range of50 to 95% and median pore diameter of 100 to 300 microns, morepreferably a median pore size between approximately 150 and 250 micronsand a porosity between 75 and 95%, which allows vascular ingrowth andthe introduction of cells into the matrix without damage to the cells orpatient. As used herein, an amorphous polymer is not crystallized; asemi-crystallized polymer is where the degree of crystallinity (fractionof mass of crystallites) is less than 100%. In general, the greater theporosity, the faster the rate of ingrowth of capillaries and connectivetissue. The rate of ingrowth is also increased by pre-wetting of thematrix with a surfactant or alcohol followed by saline wash. At thistime the most preferred embodiment is an amorphous polylactic acidhaving 90% porosity and 200 micron median pore diameter.

Polymers

Biodegradable, biocompatible polymers that degrade by hydrolysis canprovide temporary scaffolding to transplanted cells and by so doingallow the cells to secrete extracellular matrix enabling a completelynatural tissue replacement to occur. Their macromolecular structure isselected so that they are completely degraded and eliminated as the needfor an artificial support diminishes. Polymer templates for use in celltransplantation must be highly porous with large surface/volume ratiosto accommodate a large number of cells. In addition to beingbiocompatible, they must promote cell adhesion and allow retention ofdifferentiated function of attached cells. The formation of avascularized bed within the matrix for cell attachment results in anadequate supply of nutrients to transplanted cells which is essential totheir maintenance. They must also be resistant to compression and yetsemi-flexible to provide adequate support without discomfort within therecipient. Studies have been performed with poly(vinyl alcohol),although this material demonstrates some of the drawbacks of usingnon-degradable materials which may cause formation of a fibrous scar ortissue infection. Examples of useful polymers include poly(lactic acid),poly(glycolic acid), copolymers thereof, polyanhydrides,polyorthoesters, and polyphosphazines. These are all availablecommercially or can be manufactured by standard techniques.

In some embodiments, attachment of the cells to the polymer is enhancedby coating the polymers with compounds such as basement membranecomponents, agar, agarose, gelatin, gum arabic, collagens types I, II,III, IV, and V, fibronectin, laminin, glycosaminoglycans, mixturesthereof, and other materials known to those skilled in the art of cellculture.

All polymers for use in the matrix must meet the mechanical andbiochemical parameters necessary to provide adequate support for thecells with subsequent growth and proliferation. The polymers can becharacterized with respect to mechanical properties such as tensilestrength using an Instron tester, for polymer molecular weight by gelpermeation chromatography (GPC), glass transition temperature bydifferential scanning calorimetry (DSC) and bond structure by infrared(IR) spectroscopy, with respect to toxicology by initial screening testsinvolving Ames assays and in vitro teratogenicity assays, andimplantation studies in animals for immunogenicity, inflammation,release and degradation studies.

In a preferred embodiment, the matrix contains catheters for injectionof the cells into the interior of the matrix after implantation andingrowth of vascular and connective tissue. Catheters formed of medicalgrade silastic tubing of different diameters and of differing exit portsto allow even distribution of cells throughout the matrix, as describedin the following examples, are particularly useful. Other methods canalso be used, such as molding into the matrix distribution channels fromthe exterior into various parts of the interior of the matrix, or directinjection of cells through needles into interconnected pores within thematrix.

Shaping of the Matrix

The matrix is formed by methods such as casting a polymer solutioncontaining salt crystals into a mold, then leaching out the saltcrystals after the polymer is hardened, to yield a relatively rigid,non-compressible structure. This method is described in more detail inU.S. Pat. No. 5,514,378, the teachings of which are incorporated herein.

As described in more detail below, since many of the useful polymers arehydrophobic, it may be useful in some embodiments to pre-wet the matrixprior to seeding of cells within the matrix. Suitable surfactantsinclude any of the FDA approved surfactants, including polyols,alcohols, and, in some cases, saline.

Sources of Cells

In a preferred embodiment, cells are obtained either from the recipientfor autologous transplantation or from a related donor. Celltransplantation can provide an alternative treatment to whole organtransplantation for failing or malfunctioning organs such as liver andpancreas. Because many isolated cell populations can be expanded invitro using cell culture techniques, only a very small number of donorcells may be needed to prepare an implant. Consequently, the livingdonor need not sacrifice an entire organ.

Cells can also be obtained from established cell lines which exhibitnormal physiological and feedback mechanisms so that they replicate orproliferate only to a desired point.

For gene therapy, gene transfer vectors can be introduced into differentcell types, such as endothelial cells and myoblasts, which aretransplanted back to the host for the production and local release ofproteins and other therapeutic drugs. Methods for gene transfer are wellknown to those skilled in the art and have been approved by the Food andDrug Administration.

Cells types that are suitable for implantation include most epithelialand endothelial cell types, for example, parenchymal cells such ashepatocytes, pancreatic islet cells, fibroblasts, chondrocytes,osteoblasts, exocrine cells, cells of intestinal origin, bile ductcells, parathyroid cells, thyroid cells, cells of theadrenal-hypothalamic-pituitary axis, heart muscle cells, kidneyepithelial cells, kidney tubular cells, kidney basement membrane cells,nerve cells, blood vessel cells, cells forming bone and cartilage, andsmooth and skeletal muscle cells.

In one variation of the method using a single matrix for attachment ofone or more cell lines, the matrix is configured such that initial cellattachment and growth occur separately within the matrix for eachpopulation. Alternatively, a unitary scaffolding may be formed ofdifferent materials to optimize attachment of various types of cells atspecific locations. Attachment is a function of both the type of celland matrix composition. Cell attachment and viability can be assessedusing scanning electron microscopy, histology, and quantitativeassessment with radioisotopes.

Although the presently preferred embodiment is to utilize a singlematrix implanted into a host, there are situations where it may bedesirable to use more than one matrix, each implanted at the mostoptimum time for growth of the attached cells to form a functioningthree-dimensional organ structure from the different matrices.

The function of the implanted cells, both in vitro as well as in vivo,must be determined. In vivo liver function studies can be performed byplacing a cannula into the recipient's common bile duct. Bile can thenbe collected in increments. Bile pigments can be analyzed by highpressure liquid chromatography looking for underivatized tetrapyrrolesor by thin layer chromatography after being converted to azodipyrrolesby reaction with diazotized azodipyrroles ethylanthranilate either withor without treatment with P-glucuronidase. Diconjugated andmonoconjugated bilirubin can also be determined by thin layerchromatography after alkalinemethanolysis of conjugated bile pigments.In general, as the number of functioning transplanted hepatocytesincreases, the levels of conjugated bilirubin will increase. Simpleliver function tests can also be done on blood samples, such as albuminproduction. Analogous organ function studies can be conducted usingtechniques known to those skilled in the art, as required to determinethe extent of cell function after implantation. Studies using labelledglucose as well as studies using protein assays can be performed toquantitate cell mass on the polymer scaffolds. These studies of cellmass can then be correlated with cell functional studies to determinewhat the appropriate cell mass is. In most cases it is not necessary tocompletely replace the function of the organ from which the cells arederived, but only to provide supplemental or partial replacementtherapy.

Methods of Implantation

The technique described herein can be used for delivery of manydifferent cell types to achieve different tissue structures. Forexample, islet cells of the pancreas may be delivered in a similarfashion to that specifically used to implant hepatocytes, to achieveglucose regulation by appropriate secretion of insulin to cure diabetes.Other endocrine tissues can also be implanted. The matrix may beimplanted in many different areas of the body to suit a particularapplication. Sites other than the mesentery for hepatocyte injection inimplantation include subcutaneous tissue, retroperitoneum, properitonealspace, and intramuscular space.

Implantation into these sites may also be accompanied by portacavalshunting and hepatectomy, using standard surgical procedures. The needfor these additional procedures depends on the particular clinicalsituation in which hepatocyte delivery is necessary. For example, ifsignals to activate regeneration of hepatocytes are occurring in thepatient from his underlying liver disease, no hepatectomy would benecessary. Similarly, if there is significant portosystemic shuntingthrough collateral channels as part of liver disease, no portacavalshunt would be necessary to stimulate regeneration of the graft. In mostother applications, there would be no need for portacaval shunting orhepatectomy.

In the following examples, biodegradable polymer foams were prepared andimplanted into Fischer rats. The substrates utilized includepoly(L-lactic acid) and poly(DL-lactic-c-glycolic acid), which areapproved for human clinical use. Though the prevascularizationlprocedure results in the adequate supply of nutrients to attached cells,it also causes the reduction of potential space for transplanted cells.The dynamics of tissue ingrowth and vascularity, and the availabilityfor cell engraftment were determined for a variety of foams to establishtheir dependence on the polymer composition, structure and morphology.From these studies, the desired biomaterial properties, such as porosityand average pore size, were determined as well as the optimal time forcell injection.

EXAMPLE 1 Preparation of Devices

Materials and Methods

Materials

The homopolymer, poly(L-lactic acid) (PLLA), and the copolymers,poly(DL-lactic-co-glyColic) (PLGA) (85:15) and PLGA (50:50), weresupplied by Medisorb (Cincinnati, Ohio). (The ratios 85:15 and 50:50stand for the copolymer ratio of lactic acid to glycolic acid). Thepolymer molecular weights were measured by gel permeation chromatographyas M_(n)=104,800 (M_(w)/M_(n)=11.13) for PLLA, as M_(n)=121,100(M_(w)/M_(n)=1.16) for PLGA (85:15), and as M_(n)=82,800(M_(w)/M_(n)=1.14) for PLGA (50:50) (where M_(n) is the number averagemolecular weight and M_(w) is the weight average molecular weight).Granular sodium chloride (Mallinckrodt, Paris, Ky.) was ground with ananalytical mill (model A-10 Tekmar, Cincinnati, Ohio). The groundparticles were sieved with ASTM sieves placed on a sieve shaker (model18480, CSC Scientific, Fairfax, Va.). Chloroform was furnished byMallinckrodt.

Seven groups of transplantation devices were prepared by a two-stepprocedure. The devices were made of PLLA, PLGA (85:15), and PLGA(50:50). For PLLA, devices of different porosities, pore sizes, andcrystallinities were processed. First, highly porous polymer membraneswith desired porosity, pore size, and degree of crystallinity wereprepared by solvent-casting and particulate-leaching technique. Briefly,a dispersion of sieved sodium chloride particles in a chloroformsolution of PLLA (or PLGA), made by dissolution of the polymer in 8 mLof chloroform, was cast into a 5 cm Petri dish to produce a compositemembrane made of polymer and salt particles. The relative amounts ofpolymer, NaCl, and the size range of sieved NaCl particles for eachgroup are summarized in Table I. By heat treatment, the polymercrystallinity was modified, and the salt particles were leached out toyield a highly porous membrane.

The second step involved lamination of the porous membranes to constructdevices with pore structures and morphologies similar to those of theconstituent membranes. Each device was comprised of three circularlayers of diameter 13.5 mm with a medical grade silicone tubing (0.03in. inner and 0.065 in. outer diameter; American Scientific Products,McGaw Park, Ill.) inserted in the middle. At a distance of 6.75 mm fromthe stem of a knot tied at the end of a 5 cm piece of tubing, two 1/16in. rectangular holes were opened facing opposite sides for cellinjection. The thicknesses of PLLA, PLGA (85:15), and PLGA (50:50)devices were 4999 (±72), 3531 (±427), and 4484 (±296) μm, respectively.(Averages±s.d. of five measurements). For the devices made of PLLA andreported in Table I, no variation of the thickness with the porosity,pore size, or degree of crystallinity was observed. TABLE I PreparationConditions and Properties of Highly-Porous PLLA and PLGA MembranesPreparation Conditions Membrane Properties Polymer NaCl NaCl Surface/Median Pore Mass Mass NaCl Range Pore Area Volume Diameter Degree ofPolymer (g) (g) wt % (μm, μm) Porosity (cm²/mg) (1/μm) (μm)Crystallinity L90e 0.5 4.5 90 (250, 500) 0.83 1.39 0.030 166 0.245 PLLAL90c 0.5 4.5 90 (106, 150) 0.90 3.01 0.038 126 0.235 PLLA L80e 1.0 4.080 (250, 500) 0.75 1.37 0.043 137 0.245 PLLA NCL90e 0.5 4.5 90 (250,500) 0.87 1.75 0.028 179 0 PLLA NCL90c 0.5 4.5 90 (106, 150) 0.89 3.480.048 91 0 PLLA 85LG90e 0.5 4.5 90 (250, 500) 0.64 1.72 0.080 41 0 PLGA(85:15) 50LG90e 0.5 4.5 90 (250, 500) 0.84 4.93 0.106 36 0 PLGA (50:50)

The porosity, pore area, surface/volume ratio, and median pore diameterof the porous membranes were measured by mercury intrusion porosimetry,and are presented in Table I. The porosity increased as the initial saltweight fraction increased (by comparing membranes L90e and L80e) andlarger pores were formed by utilizing larger salt particles (from L90eand L90c as well as NCL90e and NCL90c). (The different codes referringto the various membranes are also included in Table I). From themeasurements of porosity and median pore diameter of the above pairs,one infers that though both properties change when either the weightpercentage of NaCl or the NaCl particle size is modified, the variationof one of them is more prominent as compared to the other. The degree ofcrystallinity of the polymers was calculated from the enthalpy ofmelting which was measured by Differential Scanning Calorimetry (7Series, Perkin-Ebmer Newton Centre, Mass.). The enthalpy of melting of100% crystallized PLLA used in the calculations was 203.4 J/g.

The devices were stored in a desiccator under vacuum until use. Theywere sterilized with ethylene oxide (12 hours exposure followed by 24hours aeration) before implantation. The sterile devices were implantedeither dry or prewet in saline just before use. The prewetting procedureinvolved dipping of devices in ethanol (100%) for one hour followed byimmersion in saline (0.9% NaCl) for at least one hour (all under sterileconditions). Prewetting of PLLA and PLGA transplantation devices wasvery important in cell seeding via injection.

EXAMPLE 2 Implantation and Harvest of Devices

Implantation into Rats

Devices were implanted in the mesentery of male 15 syngeneic Fischer 344rats (Charles River, Wilmington, Mass.). In a typical experiment, therat was anesthetized using methoxyflurane (Pitman Moore, Mundelein,Ill.) and its abdominal wall was incised. The mesentery, which is a thinfatty layer of tissue supplying blood to the small bowel, was displayedon sterile gauze carefully to avoid traumatizing the blood vessels orthe tissue. The sterile device was placed onto the unfolded mesentery.Then, the mesentery was folded back over the device so as to envelopeit, and put back into the rat. The device was sutured on the mesenteryusing non-absorbable surgical suture (Prolene, Ethicon, Sommerville,N.J.). The laparectomy was closed separately for the muscle layer andthe skin with a synthetic absorbable suture (Polyglactin 910, Ethicon).Two devices were implanted in each 200 g rat, one proximally and onedistally. For each parameter tested, a device was harvested after 5, 10,15, 20, 25, and 35 days, rinsed and stored in a 10% neutral bufferedformalin solution (Sigma, St. Louis, Mo.) until sectioning and staining.Samples were sliced into thin sections at half distance from the centerline parallel to the tubing and were stained with hematoxylin and eosin(H&E) which allowed for visualization of cells and cell nuclei.

Image Analysis

Image analysis was performed with a Megiscan 2 Image Analysis System(Joyce-Loebl, Tyne & Wear, England) equipped with a Polaroid MP-4 LandCamera. From each histological section, the area the tissue had advancedwas measured and compared to the total area of the cross section. Thearea filled by tissue was selected as the area confined between theperimeter of the harvested device and the front of the advanced tissue.A series of actual digitized sections of PLLA (with membrane code L90cas designated in Table I) devices harvested at 5, 15, 25, and 35 days isshown in FIGS. 1 a, b, c, and d, respectively. The tissue ingrowthdefined by equation (1) provided an estimate of the extent of tissueinvasion within the device.${{Tissue}\quad{Ingrowth}} = {\frac{{Area}\quad{Filled}\quad{by}\quad{Tissue}}{{Total}\quad{Area}}\left( {100\%} \right)}$

For each harvested device, the tissue ingrowth was determined from thetwo parallel and symmetric sections, and the average (±range) wascalculated. From the variation of tissue ingrowth with implantationtime, the optimal time for cell injection was determined, also referredto as prevascularization time (i.e., the time corresponding to 100%tissue ingrowth).

The regions occupied by tissue were not necessarily filled completely.From fields of 1×1 mm² outside the tissue front, the fraction of tissuearea was determined. The tissue area was measured but not the void areabecause the image of the stained polymer could not be distinguished fromthat of the vacant area. The percentage of void volume withinvascularized regions was calculated as% Void Volume=% Porosity−% Tissue Area (2)

The percentage of tissue area provides a very good estimate of thedevice volume fraction occupied by tissue. Furthermore, the frequencyand size of the invaded blood vessels were quantified. All thecapillaries within the same field were enumerated and their areadetermined. The average capillary area was calculated for each field.Three fields from the same section were used to calculate averages(±standard deviation) of the percentage of tissue area, the number ofcapillaries, and the capillary area.

For all the implants harvested, no delamination of the three-layereddevices was visually detected from the histological sections. Also, nofibrous capsule was revealed around the catheter that could havehindered the injection of cells into prevascularized devices.

Results and Discussion

PLLA Devices with 83% Porosity and 166 μm Pore Size.

The tissue ingrowth was first studied for PLLA devices of 83% porosityand 166 μm median pore diameter (L90e) which were prewet with saline,and implanted in the proximal and distal site of the mesentery for aperiod of 25 days. From FIGS. 2 a-d, one infers that the rate of tissueingrowth was reproducible and independent of the device position in themesentery. It was observed that PLLA devices 5 mm thick wereprevascularized after 25 days. The fibrovascular tissue grew into thedevices from the bases and sides. The percentage of tissue area measuredfrom histological sections increased from 67% (average of the valuesreported in FIGS. 2 a-d for devices implanted distally and proximally inthe mesentery) at day 5 to 79% at day 25. The corresponding voidfractions for cell engraftment were estimated using equation (2) as 16%at day 5 and 4% at day 25.

Effect of Prewetting

The tissue ingrowth was dependent on the prewetting of the devices.Although the same devices implanted dry were also prevascularized after25 days, as shown by FIGS. 3 a, b, c and d, the initial rate of tissueinduction was much faster for prewet devices. For example, after 5 daysof implantation, for prewet devices, the average tissue ingrowth of themeasured values for distal and proximal implantation was 44% as comparedto only 26% for the dry devices. Also, after 10 days, the relativevalues of tissue ingrowth were 74% and 44% for prewet and dry devices,respectively. Because the polymer is very hydrophobic, one infers thatprewetting reduces the adherence of the ingrowing tissue to the polymersubstrate. However, no effect was found on the total tissue area forprevascularized devices, only on the rate. The same low percentage of 4%of void volume was recorded after 25 days of implantation of drydevices. In both cases, the ingrowing tissue was highly vascularized andan appreciable increase of the tissue vascularity with the implantationtime was observed. For devices implanted dry, the average number ofcapillaries was measured as 10.0 (±5.7) at day 5 and as 17.7 (±4.6) atday 35 per field of 1×1 mm². The average capillary area was calculatedas 1100 (±610) μm² at day 5 and as 1500 (±380) μm² at day 35.

A fraction of 4% of device volume available for cell engraftment afterprevascularization is very small for an efficient transplantation ofsufficient cell mass for functional replacement. However, even if thenumber of cells fitted in the crevices between the polymer and thetissue could supplement organ function, it is very difficult to injectthat number of cells without any damage to the cells due to high shearstresses developed at their surfaces as they pass through small pores.

PLLA Devices with 75% Porosity and 137 μm Pore Size.

The dynamics of tissue ingrowth into semicrystalline PLLA devicesdepended on device porosity and pore size as deduced from the histologystudies up to 35 days, as shown by FIGS. 3 a-d. For high porosityvalues, the tissue advanced into the device much faster. After 25 days,devices of 83% porosity (L90e) were prevascularized, whereas those of75% porosity (L80e) showed a 74 (±12)% tissue ingrowth. The initialthicknesses of the devices were the same. Also, the rate of tissueingrowth was slower for devices with smaller pores.

PLLA Devices with 90% Porosity and 126 μm Pore Size.

PLLA devices of median pore diameter of 126 μm and porosity of 90%(L90c) exhibited an 85 (±1) % tissue ingrowth after 25 days. Theingrowing tissue was highly vascularized, as indicated by the measuredvalues of capillary frequency and average size over a time period of 35days. The decreased numbers of capillaries for devices of low porosity(see FIG. 3 c for L80e) do not correspond to reduced vascularity of theingrowing tissue. Rather, because the skeletal polymer volume increasesas the foam porosity decreases, they are artifacts due to the fielddefinition that also includes the space occupied by polymer. The samerationale also explains the lower values of the percentage tissue areafor the same devices.

Effect of Using Amorphous Versus Semicrystalline Devices.

Much faster tissue ingrowth occurred for amorphous PLLA devices than forsemicrystalline ones, as shown by FIGS. 4 a-d. Amorphous PLLA devices of87% porosity and median pore diameter of 179 μm (NCL90e) wereprevascularized after 20 days. Here, as for semicrystalline devices, thetissue advanced much faster into devices with larger pore diameters. Fordevices of similar porosity of 89% and median pore diameter of 91 μm(NCL90c), a tissue ingrowth of 88 (±2)% was measured at the same time.The percentage of tissue area for amorphous PLLA decreased as the foampore diameter increased. After 25 days of implantation, the ingrowingtissue filled 66 (±10)% of the NCL90e devices and 81 (±4)% of the NCL90cones, resulting in percentages of void volume for cell transplantationof 21% and 8%, respectively. The tissue vascularity was also consistent.An average number of 9.0 (±4.5) capillaries per field with an averagecross-sectional area of 845 (±593) μm² was measured for NCL90e devicesafter 25 days. This value corresponds to 13.6 capillaries per mm² oftissue, which is comparable to the number determined for the L90edevices.

The existence of sufficient space for cell engraftment renders amorphousPLLA devices of high porosity with large pores potential candidates foruse as templates for tissue regeneration. Provided that PLLA foams withdegrees of crystallinity in the range from 0% to 24.5% prepared usingthe same relative amounts of polymer and sieved salt particles havesimilar pore morphologies, as indicated from the mercury porosimetrymeasurements, one infers that the lower values of tissue area foramorphous PLLA as compared to semicrystalline PLLA reflect differentcell-polymer interactions that are not clear yet and need to beexplored.

The prevascularization of devices made of PLGA (85:15) and PLGA (50:50)was also studied. The variation of tissue ingrowth with time is shown inFIG. 5. PLGA (85:15) devices were prevascularized after 10 days, whereasfor PLGA (50:50), tissue filled the interior of the devices in 25 days.No direct comparison can be made between each other and with PLLAdevices because they not only had different pore morphologies (see TableI) but also different thickness. In addition to other possible effects,PLGA (85:15) devices were filled much faster than PLGA (50:50) onesbecause they were 953 μm thinner. The vascularity of the tissue for bothcopolymers was consistent and similar to that observed for the PLLAdevices.

In conclusion, the data demonstrate that biodegradable polymer foams ofappropriate structure and morphology can be vascularized and provide asubstrate for cell attachment and growth.

All publications and patents mentioned in this specification are hereinincorporated by reference. Although this invention has been describedwith reference to specific embodiments, variations and modifications ofthe method and means for constructing artificial organs by culturingcells on matrices having maximized surface area and exposure to thesurrounding nutrient-containing environment will be apparent to thoseskilled in the art. Such modifications and variations are intended tocome within the scope of the appended claims.

1. A polymeric matrix which can be implanted into the body, vascularizedand used as a means to achieve a high survival rate for dissociatedcells injected into the matrix, wherein the matrix is formed of abiodegradable, biocompatible, synthetic polymer, having a porositybetween 50 to 95% and a median pore size between 100 and
 300. 2. Thematrix of claim 1, wherein the pore size is between approximately 150and 250 microns and the porosity is between 75 and 95% and allowsvascular ingrowth and the introduction of cells into the matrix withoutdamage to the cells or patient.
 3. The matrix of claim 1, wherein thebiodegradable polymer is selected from the group consisting ofpolyanhydrides, polyorthoesters, polyglycolic acid, polylactic acid,copolymers and blends thereof.
 4. The matrix of claim 1, furthercomprising a wetting agent.
 5. The matrix of claim 1, wherein thepolymer is an amorphous polymer.
 6. The matrix of claim 1, wherein thepolymer is a semicrystalline polymer.
 7. The matrix of claim 1, furthercomprising means for introducing the cells into the matrix afterimplantation.
 8. The matrix of claim 1, wherein the polymer is amorphouspolylactic acid having 90% porosity and 200 micron median pore diameter.9. The matrix of claim 1, further comprising cells selected from thegroup consisting of hepatocytes, pancreatic islet cells, fibroblasts,chondrocytes, osteoblasts, exocrine cells, cells of intestinal origin,bile duct cells, parathyroid cells, thyroid cells, cells of theadrenal-hypothalamic-pituitary axis, heart muscle cells, kidneyepithelial cells, kidney tubular cells, kidney basement membrane cells,nerve cells, blood vessel cells, cells forming bone and cartilage, andsmooth and skeletal muscle cells.
 10. The matrix of claim 1, furthercomprising a material enhancing cell attachment to the polymer, whereinthe material overlays the polymer.